The present invention relates to the art of magnetic field gradient generation. It finds particular application in conjunction with establishing gradient magnetic fields in magnetic resonance imaging techniques and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in spectroscopy and other processes and apparatus in which accurately predictable magnetic field gradients are established or maintained.
In magnetic resonance imaging, a uniform magnetic field is created through an examination region in which a subject to be examined is disposed. A series of radio frequency pulses and magnetic field gradients are applied to the examination region. Gradient fields are conventionally applied as a series of gradient pulses with preselected profiles. These pulses excite magnetic resonance, phase and frequency encode the resonance, and cause phase and frequency encoded magnetic resonance signals to be emitted.
More specifically, the gradient magnetic pulses are applied to select and encode the magnetic resonance. In some embodiments, the magnetic field gradients are applied to select one or more planes or slices to be imaged. Gradient field pulses are also applied for selectively modifying the uniform magnetic field to encode frequency and phase into the magnetization, hence the resonance signals in order to identify a spatial location.
The magnetic resonance signals are then processed to generate two or three dimensional image representations of a portion of the subject in the examination region. The accuracy of the resultant image representation, among other factors, is dependent upon the accuracy with which the actually applied magnetic field gradient pulses conform to selected gradient pulse profiles.
Heretofore, linear magnetic field gradients have been produced by cylindrical gradient field coils. Discrete coils were wound in a bunched or distributed fashion on a large diameter hollow right cylinder tube, commonly 65 centimeters in diameter or larger. Conventional bunched geometries include Maxwell or modified-Maxwell pairs for z-gradient production and single or multi-arc Golay saddle coils for x and y gradient production. The coils are normally wound in a series arrangement and positioned to give a magnetic field profile with the desired linearity over a predefined volume. The distributive windings on the cylinders are generally wound and in pairs and driven anti-symmetrically. The coils are driven in an anti-symmetric manner such that only odd derivatives are non-zero at the coil origin. The first derivative is the field gradient while the third and higher order derivatives represent distortion. If the diameter of the cylinder and coil placement are chosen properly, the third derivative is canceled at the origin making the relatively weak fifth derivative component the first distortion term.
The conventional gradient coils are constructed of copper rods or multi-strand wires that are wound on a fiberglass reinforced plastic tube. The inductance, which is related to the stored magnetic energy, is critical in gradient coil design. The inductance determines to how quickly the coil can switch the gradient field on and off with a given supply of voltage. Large inductances, as are typical in wound cylindrical coils, slow the switching speed of the gradient magnetic fields.
For maximum efficiency, it would be advantageous to reduce the diameter of the gradient coil cylinders to be as close as possible to the subject, provided gradient linearity can be maintained. The required energy for field gradient production varies roughly as a fifth power of a gradient coil cylinder diameter in free space. In an actual magnetic resonance imager, the gradient coils interact with other adjoining structures such as radiation shields of superconducting magnets, making the relationship somewhat greater than the fifth power. Although reducing the coil size could have a dramatic effect on power consumption, reducing the cylinder diameter below 65 centimeters would be too restrictive to receive patients for full body scans. Although less power consumptive, a single planar surface gradient coil, by contrast, suffers from poor gradient uniformity or field linearity. Images using a single planar gradient coil require geometric distortion correction for head size volumes and larger.
The gradient coils are commonly placed on a circular cylinder for simplicity of design. The symmetry of the circular cylinder renders it relatively straightforward to calculate the gradient magnetic field at any point within the cylinder that would be caused by a series of windings on the cylinder. Generally, the winding pattern is adjusted iteratively based on mathematical calculations to obtain a linear gradient within the 65 centimeter coil. Once the designed coil becomes a physical reality, it was often necessary to make further iterative adjustments on the physical coil. These iterative adjustments for gradient uniformity are often made with little or no concern for gradient coil energy storage efficiency.
It is recognized that decreasing the volume within the coil increases coil efficiency by placing sources closer to the object examined. For a generally oval human torso, an oval gradient coil would be advantageous. However, the complexity of designing gradient coil windings on an elliptical or other non-circular cylinder is so great that it has not heretofore been achieved.
U.S. Pat. No. 4,820,988, issued Apr. 11, 1989 to Crooks, et al. shows a technique for making a more elliptical y-gradient coil. Rather than face the challenges of designing an elliptical coil, Crooks takes two circular gradient coil segments, removes the side portions adjacent their intersection with an x-z plane, and moves the remaining circular arc segments closer together. The x and z coils in Crooks were still positioned on a circular cylinder. This renders the y-gradient coil more energy efficient by reducing its patient receiving volume. It might be noted that the Crooks patent recognizes and suggests that elliptical x, y, and z coils would be advantageous. However, there is no enabling disclosure of how to construct elliptical gradient coils with sufficient linearity for magnetic resonance imager use.
In accordance with the present invention, a new and improved gradient coil configuration and method of design is provided which overcomes the above-referenced problems and others.